Coincidence counting method of γ ray and nuclear medicine diagnostic apparatus

ABSTRACT

A γ-ray signal processing section  60 ′ determines a detection time of a γ ray based on a γ-ray detection signal outputted from a semiconductor radiation detector for detecting the γ ray, and determines the energy of the γ ray. Then, a time correction circuit  70  obtains, based on the energy of the γ ray, a detection value of the detection time that corresponds to the energy of the γ ray from a time correction table indicating the relationship between the energy of the γ ray and the correction value of the detection time of the γ ray, and corrects the detection time according to the obtained correction value of the detection time. Coincidence counting is performed on the γ ray in a coincidence counting circuit  80  based on the corrected detection time.

BACKGROUND OF THE INVENTION

The present invention relates to a coincidence counting method of γ rays(gamma rays) in a nuclear medicine diagnostic apparatus, for examplepositron emission tomography (PET).

Diagnostic methods and diagnostic apparatuses are known in which aradioactive agent such as fluorodeoxyglucose (FDG) is injected to apatient and annihilation γ rays, which are emitted simultaneously inpairs in opposite directions from the body, are detected to create afunctional image (PET image). In such diagnostic methods and diagnosticapparatuses, γ rays emitted in opposite directions from the body aresimultaneously detected by two of a plurality of detectors provided inthe diagnostic apparatus, so that the incident direction (emittingdirection) of the γ ray is identified and a PET image used for a medicaldiagnosis is created. Considering that a detection time is delayed inthe detectors, processing is delayed in a detection circuit, and threeor more γ rays are not spontaneously detected, an extremely small timewindow of, e.g., 8 nsecs (8×10⁻⁹ seconds) is provided. When a subsequentγ ray is detected within 8 nsecs after the first γ ray is detected, thefirst detected γ ray and the subsequently detected γ ray are regarded asbeing emitted from the same source (γ rays emitted in a pair) and arecounted (coincidence counting), and the detection results of the γ raysare used to create a PET image. On the other hand, when a subsequent γray is not detected within 8 nsecs, a PET image is not created based onthe detection result of the first detected γ ray (JP-A-11-72566, claimsand 0021). That is, a time window with a predetermined width is set andtwo γ rays detected in the time window are judged as γ rays generated atthe same time.

γ rays are highly penetrating and hardly interact with substances.However, in some cases, γ rays interact with water and elementsconsisting of a detector and are scattered in a living body and thedetector, and the γ rays lose some energy thereof (in vivo scattering,Compton scattering). In this case, the detector does not detect γ rayswith energy of 511 keV but detects γ rays with low energy of 200 keV or400 keV, which is considerably lower than 511 keV. γ rays with lowenergy may be detected for other reasons. Moreover, annihilation γ raysof fluorine 18 (¹⁸F) has energy of 511 keV.

Conventionally, it is not possible to correctly determine the emittingdirections of the γ rays scattered with low energy (scattered ray, etc.)and thus a threshold value is set to prevent the detection results oflow energy γ rays from being used to create an image (JP-A-2003-4853,claims, 0021, FIG. 7). However, considering that valuable detectionresults are also obtained from low energy γ rays and thus sensitivityhas to be increased without placing a burden on patients and health careworkers, attempts have been made to create a PET image using scatteredray in recent years.

Coincidence counting is performed as described above also when scatteredray is used to create a PET image. Experiments conducted by the presentinventors proved a difference in detection time between a γ ray withnormal energy (γ ray in a PP area, which will be described later) and alow energy γ ray scattered by Compton scattering (γ ray in a CS area,which will be described later). When detection data with a difference indetection time on scattered ray are used, γ rays originally detected atthe same time (within a predetermined time window) may be judged as γrays detected at different times or γ rays not being detected at thesame time (within the predetermined time window) may be judged as γ raysdetected at the same time. Further, it was found that when coincidencecounting is performed over a wide range from high energy γ rays notbeing scattered (γ rays in the PP area) to low energy γ rays having beenscattered (γ rays in the CS area), the low energy γ rays cannot be useddue to the time difference unless the time window widened. Moreover, awider time window is not preferable because other γ rays with differentsources are more likely to be detected spontaneously.

Thus, it is an object of the present invention to solve the problems ofcoincidence counting using scattered ray of low energy.

SUMMARY OF THE INVENTION

A coincidence counting method of a γ ray (nuclear medicine diagnosticapparatus) of the present invention (the invention of a first view) forsolving the problems is characterized by performing (comprising) a step(function) of determining the detection time of a γ ray based on adetection signal outputted from a detector for detecting the γ ray, astep (function) of determining the energy of the γ ray based on thedetection signal of the γ ray, a step (function) of obtaining, based onthe energy of the γ ray, a correction value of the detection time from atime correction table indicating the relationship between the energy ofthe γ ray and the correction value of the detection time of the γ ray,the correction value corresponding to the energy of the γ ray, a step(function) of correcting the detection time based on the obtainedcorrection value of the detection time, and a step (function) ofperforming coincidence counting of the γ ray based on the correcteddetection time.

In the configuration according to the invention of the first aspect, thedetection time of the γ ray is corrected according to the energy of theγ ray by using the time correction table indicating the relationshipbetween the energy of the γ ray and the correction value of thedetection time of the γ ray. Besides, as will be described later, thetime correction table can be replaced with a time correction functionwhich calculates a correction value of the detection time of a γ rayaccording to the energy of the γ ray.

A coincidence counting method of a γ ray (nuclear medicine diagnosticapparatus) of the present invention (the invention of a second view) forsolving the problems is characterized by performing (comprising) a step(function) of determining the detection time of a γ ray based on adetection signal outputted from a detector for detecting the γ ray, astep (function) of determining the energy of the γ ray based on thedetection signal of the γ ray, a step (function) of calculating adifference in energy between a first detected γ ray and a subsequentlydetected γ ray which are to be subjected to coincidence counting, a step(function) of obtaining or correcting, based on the difference inenergy, a time window whose width corresponds to the difference inenergy from a time window table or time window correction tableindicating the relationship between the difference in the energy of theγ rays to be subjected to coincidence counting and the width of the timewindow, a step (function) of calculating a difference in detection timebetween the first detected γ ray and the subsequently detected γ ray,and a step (function) of comparing the time window having the obtainedor corrected width with the difference in detection time, judgingcoincidence when the difference in detection time is smaller, andperforming coincidence counting.

In the configuration according to the invention of the second aspect,the width of the time table is determined (the width of the time windowis made variable/the width of the time window is corrected) based on adifference in the energy of γ rays by using the time window table/timewindow correction table indicating the relationship between thedifference in the energy of γ rays to be subjected to coincidencecounting and the width of the time window. Further, the time windowtable/time window correction table can be also replaced with a timewindow calculating function (time window correction function).

The present invention can achieve remarkable effects as described below.

According to the coincidence counting method of a γ ray and the nuclearmedicine diagnostic apparatus of the present invention (the invention ofthe first aspect), the detection time of a γ ray is corrected accordingto the energy of the detected γ ray. Thus, even the time window has thesame width, coincidence counting can be precisely performed over a widerrange of energy. Further, for example, the time window can have a fixednarrow width, thereby minimizing the influence of other γ rays detectedspontaneously.

Further, according to the coincidence counting method of a γ ray and thenuclear medicine diagnostic apparatus of the present invention (theinvention of the second aspect), the width of the time window can bevaried (corrected) according to the energy of the detected γ ray. Thus,the coincidence counting can be precisely performed on γ rays over awide range of energy. Moreover, it is possible to minimize the influenceof other γ rays detected spontaneously as compared with the case wherethe time window remains wide.

Other objects, features and advantages of the invention will becomeapparent from the following description of the embodiments of theinvention taken in conjunction with the accompanying drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a block diagram showing an experimental apparatus used inEmbodiment 1;

FIG. 2 is a diagram showing the detecting characteristics of asemiconductor radiation detector;

FIG. 3 is a diagram showing the relationship between a detection timeand the number of detections (the number of counts) based on adifference in the energy of detected γ rays;

FIG. 4 is a diagram showing the relationship between the energy of thedetected γ ray and a peak shift (a correction value of a detection time)(a diagram showing a time correction table of Embodiment 1);

FIG. 5 is a block diagram showing the experimental apparatus used inEmbodiment 1 (with a time correction circuit);

FIG. 6 is a diagram showing experiment results obtained in theexperimental apparatus of FIG. 5;

FIG. 7 is a diagram showing the appearance of a PET apparatus accordingto Embodiment 1;

FIG. 8 is a diagram schematically showing the position relationshipbetween a patient and the semiconductor radiation detectors of the PETapparatus shown in FIG. 7;

FIG. 9 is a block diagram showing a process from the detection of a γray to the reconstruction of an image in the PET apparatus of Embodiment1;

FIG. 10 is a diagram showing a time window correction table ofEmbodiment 2;

FIG. 11 is a diagram showing a process from the detection of a γ ray tothe reconstruction of an image in the PET apparatus of Embodiment 2;

FIGS. 12A and 12B are diagrams for explaining methods of determining thedetection time of a γ ray, FIG. 12A schematically shows the principle ofa LET method; and FIG. 12B schematically shows the principle of a CFDmethod.

DESCRIPTION OF THE EMBODIMENTS

Preferred embodiments (hereinafter abbreviated as “embodiments”) forimplementing a coincidence counting method of γ rays and a nuclearmedicine diagnostic apparatus will be specifically described below withreference to the accompanying drawings.

Embodiments 1 and 2 described below use scattered ray. Embodiment 1 hasa fixed time window and Embodiment 2 has a variable time window.

Embodiment 1

First, regarding Embodiment 1 using scattered ray with a fixed timewindow, a principle, a concrete example, and the operations of theconcrete example will be discussed in this order.

[Principle]

The principle of Embodiment 1 will be discussed below.

First, the present inventors clarified the relationship between energy(peak value) of a γ ray detected by a semiconductor radiation detector51 and a difference in detection time (detection time difference) byusing an experimental apparatus simplified in FIG. 1. The experimentalapparatus of FIG. 1 comprises the semiconductor radiation detector 51which acts as a detector for detecting γ rays and is made of cadmiumtelluride (CdTe), a circuit (γ-ray signal processing section 60′) whichhas the function of determining the energy of a detected γ ray anddetermining a detection time, and a coincidence counting circuit 80which has the function of measuring a detection time difference relativeto energy based on the energy and the detection time.

Annihilation γ rays (γ rays in a pair emitted in opposite directions by180°) of ¹⁸F with 511 keV that are generally used in PET examinationswere generated in the experimental apparatus, and experiments wererepeated to measure the energy of γ rays detected by the semiconductorradiation detector 51 and a difference in detection time of the γ rays.

As a result, graphs shown in FIGS. 2 and 3 were obtained. In FIG. 2, ahorizontal axis represents the energy of γ rays and a vertical axisrepresents the number of detections of γ rays (the number of counts) InFIG. 3, a horizontal axis represents a detection time difference and avertical axis represents the number of detections of γ rays (the numberof counts) In FIG. 2, an area from 450 to 550 keV including 511 keV is aphotopeak area (PP). Further, an area from 200 (or 150) to 550 keV is aphotopeak (PP) area+Compton scattering (CS) area. Incidentally, in theCS area, γ rays originally having energy of 511 keV are actuallydetected with energy lower than that of the PP area due to Comptonscattering.

As indicated by a broken line in FIG. 3, a difference in detection timeof γ rays in the PP area of FIG. 2 is within about ±8 nsec. That is,regarding the energy of a pair of γ rays detected in the PP area, it ispossible to positively perform coincidence counting on the γ raysgenerated in a pair by setting a time window with a width of 8 nsec. Adifference in detection time between γ rays in the PP area+CS area inFIG. 2 is within about ±12 nsec (16 nsec) as indicated by a solid linein FIG. 3. That is, it is understood that regarding the energy of thepair of detected γ rays in the PP area+CS area, in other words, when γrays down to 200 keV are subjected to coincidence counting, coincidencecounting can be performed almost positively on γ rays generated in apair by setting a time window with a width of 12 nsec (16 nsec).

However, it is also understood that simply with a wider time window, itis more likely to spontaneously detect other γ rays as γ rays detectedat the same time (simply a wider time window reduces sensitivity ineffect).

It was found that as shown in FIG. 3 by dashed lines, when one of γ raysgenerated in a pair is present in the PP area of 500 keV (γ ray detectedin the PP area) and the other is present in the CS area of 300 keV (γray detected in the CS area), a detection time difference of 4 nsec ismade at the median value. In other words, it was found that the peak ofthe detection time difference is shifted by 4 nsec. The presentinventors further examined the facts and found that as detected γ rayshave lower energy, detection time is delayed (FIG. 4). Further, theinventors found that coincidence counting using γ rays (γ rays from200(150) to 550 keV) widely from the PP area to the CS area can beperformed by correcting the delay without widening the time window.

In FIG. 4 showing an example of a time correction table, a horizontalaxis represents the energy of detected γ rays (keV) and a vertical axisrepresents a peak shift (nsec) serving as a correction value of adetection time. In FIG. 4, a peak shift of a γ ray is plotted at eachenergy with 511 keV (511 keV or more) serving as a peak shift of 0. Itis understood that as energy decreases, a peak shift increases(increases in + direction). That is, FIG. 4 shows that when a γ ray withenergy 400 keV is detected, the detection time (measurement value) isdelayed by 2 nsec relative to a γ ray with energy of 511 keV (peakshift+2 nsec). Further, FIG. 4 shows that when a γ ray with energy of300 keV is detected, the detection time (measurement value) is delayedby 4 nsec relative to a γ ray with energy of 511 keV (peak shift+4 nsec)Further, FIG. 4 shows that when a γ ray with energy of 200 keV isdetected, the detection time (measurement value) is delayed by 8 nsecrelative to a γ ray with energy of 511 keV (peak shift+8 nsec).Therefore, the following is understood: when a γ ray with energy of 400keV is detected, the detection time is advanced by 2 nsec (thecorrection value of a peak shift−2 nsec), when a γ ray with energy of300 keV is detected, the detection time is advanced by 4 nsec (acorrection value of a peak shift−4 nsec), and when a γ ray with energyof 200 keV is detected, the detection time is advanced by 8 nsec(correction value of the peak shift−8 nsec).

Referring to FIG. 4, a concrete example of peak correction will bediscussed below. When successively detected γ rays have energy of 511keV, neither of the γ rays requires peak shift correction. Further, whenone of successively detected γ rays has energy of 511 keV and the otherhas energy of 400 keV, peak shift correction is performed to subtract 2nsec from the detection time of the γ ray of 400 keV. When one ofsuccessively detected γ rays has energy of 511 keV and the other hasenergy of 200 keV, peak shift correction is performed to subtract 8 nsecfrom the detection time of the γ ray of 200 keV.

Moreover, when one of successively detected γ rays has energy of 400 keVand the other also has energy of 400 keV, peak shift correction isperformed to subtract 2 nsec from the detection time of each of the γrays. When one of successively detected γ rays has energy of 400 keV andthe other has energy of 300 keV, peak shift correction is performed tosubtract 2 nsec from the detection time of the γ ray of 400 keV and peakshift correction is performed to subtract 4 nsec from the detection timeof the γ ray of 300 keV.

That is, a peak shift is corrected (detection time is corrected)according to the energy of a detected γ ray.

When the energy of a γ ray is a median value including 260 keV and 385keV, a peak shift (correction value) can be determined from energyvalues by using a time correction table and a time correction map whichdescribe more specific values, and correlation equations (timecorrection functions) which include a linear function, a secondaryfunction, and an exponential function and indicate the relationshipbetween energy and a peak shift. That is, even when the time correctiontable is replaced with a time correction function for calculating acorrection value of the detection time of a γ ray according to theenergy of the detected γ ray, the object of the present invention can beattained. For this reason, it is assumed that the time correctionfunction is included in the concept of the time correction table (timecorrection table=time correction function).

Incidentally, FIG. 4 shows that 511 keV is set at a peak shift of 0 asan example. It is needless to say that 400 keV or 200 keV may be set at0.

Subsequently, the present inventors conducted experiments for correctingthe detection time of a γ ray according to the energy of the γ raydetected by the semiconductor radiation detector 51, by using anexperimental apparatus simplified in FIG. 5. The experimental apparatusof FIG. 5 is different from that of FIG. 1 in that a time correctioncircuit 70 and a time correction table used in the time correctioncircuit 70 are provided in the previous stage of a coincidence countingcircuit 80. Other configurations are similar to those of theexperimental apparatus shown in FIG. 1 and thus the explanation thereofis omitted. A method of correcting a detection time is also similar tothe above described method, in which a corresponding correction value isobtained from the time correction table according to the energy of adetected γ ray and the correction value is subtracted from the detectiontime. Thus, further explanation is omitted.

FIG. 6 shows experimental results obtained in the experimental apparatusof FIG. 5. In FIG. 6, a relatively thick solid line represents a curveafter correction and a relatively thin solid line represents a curvebefore correction (similar to the curve of the solid line of FIG. 3). Asshown in FIG. 6, a detection time difference of ±12 nsec (16 nsec)before correction could be reduced to ±8 nsec. Accordingly, the peak ofthe curve could be increased. That is, FIG. 6 shows that a time windowof 8 nsec enables proper coincidence counting to be performed on γ raysover a wide range from 200 to 550 keV (PP area+CS area). Further, FIG. 6shows that the influence of other γ rays detected spontaneously can beminimized by setting such a narrow time window.

[Concrete Example]

Referring to FIG. 7 and others, a concrete example of Embodiment 1 willbe discussed below.

As shown in FIGS. 7 and 8, a PET apparatus 1 serving as a nuclearmedicine diagnostic apparatus includes a camera (imaging device) 11, adata processor 12, and a display device 13. An examinee (patient) islaid on a bed 14 and is photographed by the camera 11. The camera 11comprises a number of detectors (semiconductor radiation detectors) fordetecting γ rays. The detectors detect γ rays emitted from the body ofthe examinee on the bed 14. The camera 11 has an integrated circuit(ASIC) for measuring a peak value and a detection time of a γ ray. Theintegrated circuit measures a peak value (energy) and a detection timeof a detected γ ray. The data processor 12 has a coincidence countingcircuit 80, a storage device 90, and an image reconstructing device 100(FIG. 8). The data processor 12 captures a peak value of a detected γray, the data of a detection time, and packet data including a detector(channel) ID. The coincidence counting circuit 80 performs coincidencecounting based on the packet data, particularly the data of thedetection time and the detector ID (address N), locates the detectionposition of a γ ray of 511 keV, and stores the position in the storagedevice 90. The image reconstructing device 100 creates a functionalimage based on information stored in the storage device 90 and displaysthe image on the display device 13.

The semiconductor radiation detectors serving as detectors are arrangedlike a circle in the camera 11 so as to surround the patient.Incidentally, the patient is administered with a radioactive agent,e.g., fluorodeoxyglucose (FDG) including 18F having a half-life of 110minutes. γ rays (annihilation γ rays) are emitted from the body of theexaminee when positrons emitted from the FDG are annihilated.

As shown in FIG. 9, the PET apparatus 1 includes a γ-ray detectionsignal processing section 60, the coincidence counting circuit 80, thestorage device 90, and the image reconstructing device 100. The γ-raydetection signal processing section 60 is provided in the camera 11. Thecoincidence counting circuit 80, the storage device 90, and the imagereconstructing device 100 are provided in the data processor 12. Theγ-ray detection signal processing section 60 and the data processor 12(the coincidence counting circuit 80, the storage device 90, and theimage reconstructing device 100) correspond to “processors.” AlthoughFIG. 9 shows only one (one system) γ-ray detection signal processingsection 60, several tens (several tens systems) or more γ-ray detectionsignal processing sections 60 are actually provided. Thus, thecoincidence counting circuit 80 comprises signal input ports as many asthe γ-ray detection signal processing sections 60.

Reference numeral 51 (51 a, 51 b, . . . 51 n) denotes a semiconductorradiation detector. In FIG. 9, subscripts a, b, . . . n of referencenumerals 51, 61, and others are used when configurations areindividually explained, and the subscripts are omitted in other cases.

The γ-ray detection signal processing section 60 of FIG. 9 processes aγ-ray detection signal outputted when the semiconductor radiationdetector 51 detects a γ ray. The γ-ray detection signal processingsection 60 includes a preamplifier circuit 61, a low-speed waveformamplifier circuit 62, a γ-ray discriminator circuit 63, a pulse-heightanalyzer circuit 64, a high-speed amplifier circuit 65, a time pick-offcircuit 66, a high-speed clock 67, an address discriminating circuit 68,an event data output circuit 69, and a time correction circuit 70. Thelow-speed waveform amplifier circuit 62, the γ-ray discriminator circuit63, and the pulse-height analyzer circuit 64 constitute a low-speedsignal processing system which determines the energy (E) of a detected γray. The high-speed amplifier circuit 65, the time pick-off circuit 66,the high-speed clock 67, and the address discriminating circuit 68constitute a high-speed signal processing system which determines adetection time (τ) and an address of a γ ray (N→the ID of thesemiconductor radiation detector which detects the γ ray). Thepulse-height analyzer circuit 64 and the address discriminating circuit68 are fed with signals from the semiconductor radiation detectors 51 a,51 b, . . . 51 n. In this case, the high-speed clock 67 generates aclock with a speed for determining a detection time in nsec. Thepreamplifier circuit 61 and so on are provided, as indicated byreference numerals 61 a, 61 b, . . . 61 n, in the γ-ray detection signalprocessing section 60 as many as the semiconductor radiation detectors51 a, 51 b, . . . 51 n.

The semiconductor radiation detector 51 includes a semiconductor element52, an anode 53, and a cathode 54. A voltage of several hundreds voltsfor collecting charge is applied between the anode 53 and the cathode 54of the semiconductor radiation detector 51 by a voltage supply 55.

The configurations indicated by reference numerals 61 to 69 of they-raydetection signal processing section 60 have been already clarified by avariety of documents, and thus the detailed explanation thereof isomitted. These configurations may be replaced with other configurationsdifferent from those of FIG. 9 as long as the detection time (τ) of a γray and the energy (E) of a detected γ ray are outputted to the timecorrection circuit 70.

As shown in FIG. 9, the γ-ray detection signal processing section 60 ofthe present embodiment comprises the time correction circuit 70. Thetime correction circuit 70 comprises the time correction table (notshown in FIG. 9, see FIG. 4) or time correction function and has thefunction of setting a correction value (τ_(E)) of a detection timeaccording to the energy of a γ ray. Further, the time correction circuit70 has the function of correcting a detection time (τ-τ_(E)) by using acorrection value. Incidentally, in the present embodiment, γ rays withenergy of 200 keV or higher are subjected to coincidence counting. γrays less than 200 keV are excluded out of the series of processing by,e.g., the γ-ray discriminator circuit 63 and the pulse-height analyzercircuit 64 (cut by a threshold value). γ rays less than 150 keV may beexcluded out of the series of processing and γ rays of 150 keV or highermay be subjected to coincidence counting. Under normal conditions, γrays of less than 450 keV are excluded out of the series of processing.

The address (N), the corrected detection time (τ-τ_(E)), and the energy(E) are outputted from the time correction circuit 70 to the coincidencecounting circuit 80 of the data processor 12.

The coincidence counting circuit 80 has the function of coincidencecounting using the first inputted signal (signals (N, τ-τ_(E), E) fromthe γ-ray detection signal processing section 60 of one system) and thesubsequently inputted signal (signals (N, τ-τ_(E), E) from the γ-raydetection signal processing section 60 of another system). To bespecific, the coincidence counting circuit 80 has the function ofdetermining a detection time difference when a signal is first inputtedfrom the γ-ray detection signal processing section 60 of one system anda signal is subsequently inputted from the γ-ray detection signalprocessing section 60 of another system, the function of decidingwhether a detection time difference is within a time window (e.g., 8nsec), the function of outputting the inputted address (N) of one systemand the address (N) of another system to the storage device 90 in thesubsequent stage when the detection time difference is within the timewindow, and the function of excluding the inputted signal when thedetection time difference exceeds the time window.

The configuration of the coincidence counting circuit 80 has beenalready clarified by a variety of documents, and the configuration maybe replaced with another configuration. Further, the width of the timewindow can be set as appropriate.

The storage device 90 stores the address (N) of the semiconductorradiation detector that has been subjected to coincidence counting. Theimage reconstructing device 100 reads the contents stored in the storagedevice 90, determines the emitting direction of a γ ray, andreconstructs a PET image. The storage device 90 and the imagereconstructing device 100 are similar to conventional ones and thus theexplanation thereof is omitted.

[Operations of the Concrete Example]

The following will discuss operations of the PET apparatus 1 configuredas the above described concrete example (See FIGS. 4 to 9 whennecessary).

When a pair of γ rays emitted from the body of the examinee are detectedby the semiconductor radiation detectors 51 (the semiconductor radiationdetector 51 of one system and the semiconductor radiation detector 51 ofanother system), each of the semiconductor radiation detectors 51outputs a γ-ray detection signal. Regarding the outputted γ-raydetection signals, the energy of the γ ray is determined by a slowprocessing system of the γ-ray detection signal processing section 60 inthe corresponding system, and a detection time and the address of thesemiconductor radiation detector 51 are determined by a fast processingsystem. The detection time is corrected by the time correction circuit70 according to the energy of the γ ray. The signal including thecorrected detection time is inputted to the coincidence counting circuit80. Then, the coincidence counting circuit 80, which have been fed withthe signal from the γ-ray detection signal processing section 60 and thesignal from the γ-ray detection signal processing section 60 of anothersystem, determines a detection time difference based on the correcteddetection time. When the detection time difference is within the timewindow, the addresses (N) of the semiconductor radiation detectors 51having detected the pair of γ rays are outputted to the storage device90. Then, the image reconstructing device 100 reconstructs a PET imagebased on data of the addresses (N) stored in the storage device 90, anddisplays the PET image on the display device 13.

According to Embodiment 1, a γ-ray detection signal (scattered ray) inthe CS area is also used for the PET image displayed on the displaydevice 13, thereby achieving higher sensitivity than the conventionalart. Additionally, since the γ-ray detection signal in the CS area isused without expanding the time window, other γ rays are less likely tobe detected spontaneously (no reduction in sensitivity). Thus, it ispossible to shorten examining time and to reduce a dose of a radioactiveagent administered to the examinee. The shorter examining time, thenumber of times of processing (the number of examinations) can beincreased. Further, a smaller dose of the radioactive agent administeredto the examinee results in the smaller exposure of the examinee andmedical workers.

Incidentally, when none of γ rays emitted in a pair from the body of theexaminee is detected, as a matter of course, any image is notreconstructed from the γ rays. Moreover, also when only one of γ rays ina pair is detected, any image is not reconstructed from the γ rays(except for spontaneous coincidence counting) in consideration of thetime window of 8 nsec. According to Embodiment 1, by correcting adetection time, the time window can be set at a short time period whileusing scattered ray. Hence, as described above, other γ rays are lesslikely to be detected spontaneously.

An explanation about processing for scattered ray is omitted. Theprocessing for scattered ray is performed according to, e.g., Comptonscattering. For example, when energy of 200 keV is applied to the firstsemiconductor radiation detector 51 and the remainder of the energy isapplied to an adjacent semiconductor radiation detector 51, both of γrays are regarded as a single γ ray of scattered ray in view of therelationship among a total value of energy of the γ rays, a detectiontime, and the addresses of the semiconductor radiation detectors 51. Aconfiguration for the scattered ray processing may be provided in, e.g.,the event data output circuit 69 and the coincidence counting circuit 80(in the previous stage of the coincidence counting circuit 80).Incidentally, when a detection time is corrected after the detectiontime is determined for a γ ray having been subjected to the scatteredray processing, the current configuration for the scattered rayprocessing is used as it is with enhanced convenience. In this case, thedetection time can be corrected according to the energy and detectiontime of the first detected γ ray.

As a matter of course, scattered ray can be properly used without theneed for a wider time window or the configuration for scattered rayprocessing.

Incidentally, a peak shift of FIG. 4 (a correction value of a detectiontime) is varied for each of the semiconductor radiation detectors 51 andthe γ-ray detection signal processing sections 60, and thus statisticsmay be organized to absorb the individual differences.

Embodiment 2

Regarding Embodiment 2 using scattered ray with a variable time window,the following will discuss a principle, a concrete example, and theoperations of the concrete example in this order. An explanation aboutthe same parts as Embodiment 1 is omitted.

[Principle]

The principle of Embodiment 2 will be discussed below.

According to the experiments using the experimental apparatus (FIG. 1)described in Embodiment 1, the present inventors found that a differencein detection time (detection time difference) occurs according to theenergy of a γ ray (FIG. 3). Based on this knowledge, a detection time iscorrected in Embodiment 1 (FIGS. 4 and 5). In Embodiment 2, a timewindow is corrected instead of a detection time.

In FIG. 10 showing an example of a time window correction table, ahorizontal axis represents a difference in energy between successivelydetected γ rays (keV) and a vertical axis represents a correction value(nsec) of the time window. As is understood from FIG. 10, in the case ofno difference in energy between successively detected γ rays, the timewindow has a correction value of 0 nsec. For example, when the timewindow is set at 8 nsec (no difference in energy, e.g., 511 keV-511 keVand 200 keV-200 keV), the time window remains 8 nsec. When a differencein energy is 300 keV (200 keV-511 keV) between successively detected γrays, the time window has a correction value of 8 nsec. For example,when the time window is set at 8 nsec, the time window is increased to16 nsec.

As shown in FIG. 4, a peak shift is represented as a curve (particularlyan area of 300 keV or lower). Therefore, as shown in FIG. 10, thepresent embodiment has a time window correction table according to theenergy of the first detected γ ray (time window correction map/timewindow correction function) for convenience. Incidentally, when a valueof the time window correction table of FIG. 10 is offset above by 8nsec, the value of the time window can be obtained based on a differencein energy in the time window table. Although FIG. 10 shows two lines of200 keV and 511 keV, it is preferable to provide more lines of 300 keV,400 keV, and so on to minutely correct the time window.

[Concrete Example]

Referring to FIG. 11 and others, a concrete example of Embodiment 2 willbe discussed below.

A PET apparatus 1′ of Embodiment 2 is similar in appearance to that ofEmbodiment 1 (FIG. 7). As shown in FIG. 11, a γ-ray detection signaldetected by a semiconductor radiation detector 51 is processed in aγ-ray detection signal processing section 60′. Unlike the γ-raydetection signal processing section 60 of Embodiment 1, the γ-raydetection signal processing section 60′ has a typical configuration nothaving a time correction circuit 70 (or a time correction table). Thisis because Embodiment 2 makes a time window variable instead ofcorrecting the detection time of a γ ray.

As shown in FIG. 11, a coincidence counting circuit 80′ of Embodiment 2comprises a time window correction table 81. The coincidence countingcircuit 80′ has the function of making the time window variable andperforming coincidence counting using the first inputted signal (signals(N, τ, E) from the γ-ray detection signal processing section 60′ of onesystem) and the subsequently inputted signal (signals (N, τ, E) from theγ-ray detection signal processing section 60′ of another system).

To be specific, the coincidence counting circuit 80′ has the function ofdetermining a difference in energy between signals when a signal isfirst inputted from the γ-ray detection signal processing section 60′ ofone system and a signal is subsequently inputted from the γ-raydetection signal processing section 60′ of another system (when signalsare successively inputted), the function of obtaining a correction valueof the time window, the correction value corresponding to the energydifference from the time correction table according to the energydifference, the function of correcting the time window according to thecorrection value of the time window and making the time window variable,the function of determining a detection time difference between γ raysof the systems, the function of deciding whether the detection timedifference is within the variable time window, the function ofoutputting the inputted address (N) of one system and the address (N) ofanother system to a storage device 90 in the subsequent stage when thedetection time difference is within the time window, and the function ofexcluding the inputted signal when the detection time difference exceedsthe time window.

The storage device 90 and an image reconstructing device 100 are similarto those of Embodiment 1 and thus the explanation thereof is omitted.

[Operations of the Concrete Example]

The following will discuss the operations of the PET apparatus 1′configured as the above described concrete example (See FIGS. 10 and 11when necessary).

When a pair of γ rays emitted from the body of an examinee are detectedby the semiconductor radiation detectors 51 (the semiconductor radiationdetector 51 of one system and the semiconductor radiation detector 51 ofanother system), each of the semiconductor radiation detectors 51outputs a γ-ray detection signal. Regarding the outputted γ-raydetection signal, the energy (E) of the γ ray is determined by a slowprocessing system of the γ-ray detection signal processing section 60′in the corresponding system, and the detection time (τ) and the address(N) of the semiconductor radiation detector 51 are determined by a fastprocessing system.

Then, the coincidence counting circuit 80′, which have been successivelyfed with the signal (N, τ, E) from the γ-ray detection signal processingsection 60′ and the signal (N, τ, E) from the γ-ray detection signalprocessing section 60 of another system, determines a difference inenergy, obtains a correction time of the corresponding time window fromthe time window correction table, and corrects the time window (makesvariable). Meanwhile, a detection time difference is determined. Whenthe detection time difference is within the variable time window, theaddresses (N) of the semiconductor radiation detectors 51 havingdetected the pair of γ rays are outputted to the storage device 90.Then, the image reconstructing device 100 reconstructs a PET image basedon data of the addresses (N) stored in the storage device 90, anddisplays the PET image on the display device 13.

According to Embodiment 2, a γ-ray detection signal (scattered ray) ofthe CS area is also used for the PET image displayed on the displaydevice 13, thereby achieving higher sensitivity than the conventionalart. Additionally, the width of the time window is made variableaccording to the difference in the energy of γ rays, and thus the γ-raydetection signal can be used with a small increase in the time window.Hence, other γ rays are less likely to be spontaneously detected (asmaller reduction in sensitivity). Thus, it is possible to shortenexamining time of a camera 11 (FIG. 7) and to reduce a dose of aradioactive agent administered to an examinee. As the examining timedecreases, the number of times of processing (the number ofexaminations) per hour can be increased for the examinee. Further, asthe dose of the radioactive agent administered to the examineedecreases, the exposure of the examinee and medical workers can bereduced.

Scattering radiation processing is similar to that of Embodiment 1. Adifference in energy is varied for each of the semiconductor radiationdetectors 51 and the γ-ray detection signal processing sections 60′, andthus statistics may be organized to absorb the individual differences asin Embodiment 1.

<Supplementary Explanation>

A supplementary explanation will be made on a method of determining thedetection time of a γ ray.

The method of determining the detection time of a γ ray is broadlyclassified as Leading Edge Trigger (LET) whose principle isschematically shown in FIG. 12A and Constant Fraction Discriminator(CFD) whose principle is schematically shown in FIG. 12B.

As shown in FIG. 12A, the LET is a method of determining the detectiontime of a γ ray, in which a detection time is determined when the energy(intensity) of a γ ray rises and reaches a set threshold value (Vth).According to the LET method, a detection time can be swiftly determinedwith ease, whereas the detection time of a low energy γ ray tends to bedelayed as compared with a high energy γ ray.

FIG. 12B shows time-energy (intensity) curves of γ rays according to theCFD method. A curve obtained by multiplying measurement values by, e.g.,0.2 times and a curve with unchanged measurement values are shifted inthe horizontal axis (time base) direction, the curves are shifted alsoin the vertical axis direction (add Walk), and the intersection of thecurves is determined as the detection time of the γ ray. According tothe CFD method, a difference in detection time can be small from lowenergy γ rays to high energy γ rays. However, a difference in detectiontime still occurs because of Walk (offset voltage).

The present invention is applicable to both of the methods, andparticularly the LET method is effective. Incidentally, a detection timeof a γ ray is determined by the LET method in Embodiment 1 andEmbodiment 2. The operations of the LET and CFD methods are performed ina time pick-off circuit 66 shown in FIG. 9.

The present invention described above is not limited to Embodiment 1 andEmbodiment 2. Variations may be made within the scope of the technicalidea.

For example, when coincidence counting is performed only by using γ raysin the PP area, a time window may be narrowed by using theconfigurations of Embodiment 1 and Embodiment 2.

Further, the present invention is also applicable to a radiationdetector such as a scintillator other than the semiconductor radiationdetector.

It should be further understood by those skilled in the art thatalthough the foregoing description has been made on embodiments of theinvention, the invention is not limited thereto and various changes andmodifications may be made without departing from the spirit of theinvention and the scope of the appended claims.

1. A method of determining a detection time of γ ray in a nuclear medicine diagnostic apparatus constructing a functional image by processing γ ray detection signals output from a plurality of sensors for detecting γ rays, comprising the steps of: determining energy of γ rays on the basis of said γ ray detection signals output from said sensors; and determining a γ ray detection time on the basis of said γ ray detection signals and said energy of γ rays.
 2. A method of determining a detection time of γ ray in a nuclear medicine diagnostic apparatus constructing a functional image by processing γ ray detection signals output from a plurality of sensors for detecting γ rays, comprising the steps of: measuring a γ ray detection time on the basis of said γ ray detection signals output from said sensors; determining energy of γ rays on the basis of said γ ray detection signals; and correcting said γ ray detection time measured on the basis of said energy of γ rays.
 3. A method of determining a detection time of γ ray in a nuclear medicine diagnostic apparatus constructing a functional image by processing γ ray detection signals output from a plurality of sensors for detecting γ rays, comprising the steps of: measuring a γ ray detection time on the basis of said γ ray detection signals output from said sensors; determining energy of γ rays on the basis of said γ ray detection signals; and correcting said γ ray detection time measured using a correction value of a detection time corresponding to said energy of γ rays.
 4. A coincidence counting method of γ rays in a nuclear medicine diagnostic apparatus constructing a functional image by processing γ ray detection signals output from a plurality of sensors for detecting γ rays, comprising the steps of: determining energy of γ rays on the basis of said γ ray detection signals output from said sensors; determining a γ ray detection time on the basis of said γ ray detection signals and said energy of γ rays; and executing coincidence counting of γ rays on the basis of said detection time.
 5. A coincidence counting method of γ rays in a nuclear medicine diagnostic apparatus constructing a functional image by processing γ ray detection signals output from a plurality of sensors for detecting γ rays, comprising the steps of measuring a γ ray detection time on the basis of said γ ray detection signals output from said sensors; determining energy of γ rays on the basis of said γ ray detection signals; correcting γ ray detection time measured on the basis of said energy of γ rays; and executing coincidence counting of γ rays on the basis of a detection time obtained in said correcting step.
 6. The coincidence counting method of γ rays according to claim 5, wherein said step of executing coincidence counting comprises a step of judging said γ rays as being coincident, when a difference between detection times obtained in said step of correcting is smaller than a predetermined time window.
 7. The coincidence counting method of γ rays according to claim 5, wherein said step of executing coincidence counting comprises a step of deciding whether or not said γ rays detected should be subjected to coincidence counting as scattered radiations on the basis of γ ray detection time obtained in said correcting step and said determined energy.
 8. The coincidence counting method of γ rays according to claim 5, wherein said step of executing coincidence counting comprises a step of removing said γ rays from γ rays to be subjected to coincidence counting, when said energy of said γ rays determined is below a predetermined threshold value.
 9. A coincidence counting method of γ rays in a nuclear medicine diagnostic apparatus constructing a functional image by processing γ ray detection signals output from a plurality of sensors for detecting γ rays, comprising the steps of: measuring a γ ray detection time on the basis of said γ ray detection signals output from said sensors; determining energy of γ rays on the basis of said γ ray detection signals; correcting said γ ray detection time measured using a correction value of a detection time corresponding to said energy of γ rays; and executing coincidence counting of γ rays on the basis of a detection time obtained in said correcting step.
 10. The coincidence counting method of γ rays according to claim 9, wherein said correction value and said energy of γ rays are correlated in a manner that the smaller said energy of said γ rays, the larger said correction value to be subtracted from said detection time.
 11. A nuclear medicine diagnosis apparatus comprising: a plurality of sensors for detecting γ rays; an energy determination unit for determining energy of said γ rays on the basis of γ ray detection signals output from said sensors; a time determination unit for determining a γ ray detection time on the basis of said γ ray detection signals and said energy of said γ rays; and a coincidence counting unit for executing coincidence counting of said γ rays on the basis of said γ ray detection time.
 12. The nuclear medicine diagnosis apparatus according to claim 11, further comprising a decision unit for deciding as to whether or not said γ rays detected should be subjected to coincidence counting on the basis of said γ ray detection time and said energy of γ rays.
 13. The nuclear medicine diagnosis apparatus according to claim 11, wherein said γ rays are removed from γ rays to be subjected to coincidence counting, when said energy of γ rays is below a predetermined threshold value.
 14. A nuclear medicine diagnosis apparatus comprising: a plurality of sensors for detecting γ rays; a detection time measurement unit for measuring a γ ray detection time on the basis of γ ray detection signals output from said sensors; an energy determination unit for determining energy of γ rays on the basis of said γ ray detection signals; a detection time correction unit for correcting said γ ray detection time on the basis of said energy of γ rays; and a coincidence counting unit for executing a coincidence counting of γ rays on the basis of a γ ray detection time obtained in said means for correcting.
 15. The nuclear medicine diagnosis apparatus according to claim 14, wherein said coincidence counting unit judges said γ rays as being coincident when a difference between γ ray detection times obtained in said detection time correction unit is smaller than a predetermined time window and executes the coincidence counting.
 16. A nuclear medicine diagnosis apparatus comprising: a plurality of sensor for detecting γ rays; a detection time measurement unit for measuring a γ ray detection time on the basis of γ ray detection signals output from said sensors; an energy determination unit for determining energy of said γ rays on the basis of said γ ray detection signals; a detection time correction unit for correcting said γ ray detection time using a correction value of a γ ray detection time corresponding to said energy of γ rays; and a coincidence counting unit for executing a coincidence counting of γ rays on the basis of a γ ray detection time obtained in said means for correcting. 